Images of the interiors of bodies may be acquired using various types of tomographic techniques, which involve recording and measuring radiation from tissues and processing acquired data into images.
One of these tomographic techniques is positron emission tomography (PET), which involves determining spatial distribution of a selected substance throughout the body and facilitates detection of changes in the concentration of that substance over time, thus allowing to determine the metabolic rates in tissue cells.
The selected substance is a radiopharmaceutical administered to the examined object (e.g. a patient) before the PET scan. The radiopharmaceutical, also referred to as an isotopic tracer, is a chemical substance having at least one atom replaced by a radioactive isotope, e.g. 11C, 15O, 13N, 18F, selected so that it undergoes radioactive decay including the emission of a positron (antielectron). The positron is emitted from the atom nucleus and penetrates into the object's tissue, where it is annihilated in reaction with an electron present within the object's body.
The phenomenon of positron and electron annihilation, constituting the principle of PET imaging, consists in converting the masses of both particles into energy emitted as annihilation photons, each having the energy of 511 keV. A single annihilation event usually leads to formation of two photons that diverge in opposite directions at the angle of 180° in accordance with the law of conservation of the momentum within the electron-positron pair's rest frame, with the straight line of photon emission being referred to as the line of response (LOR). The stream of photons generated in the above process is referred to as gamma radiation and each photon is referred to as gamma quantum to highlight the nuclear origin of this radiation. The gamma quanta are capable of penetrating matter, including tissues of living organisms, facilitating their detection at certain distance from object's body. The process of annihilation of the positron-electron pair usually occurs at a distance of several millimeters from the place of the radioactive decay of the isotopic tracer. This distance constitutes a natural limitation of the spatial resolution of PET images to a few millimeters.
A PET scanner comprises detection devices used to detect gamma radiation as well as electronic hardware and software allowing to determine the position of the positron-electron pair annihilation event on the basis of the position and time of detection of a particular pair of the gamma quanta. The radiation detectors are usually arranged in layers forming a ring around object's body and are mainly made of an inorganic scintillation material. A gamma quantum enters the scintillator, which absorbs its energy to re-emit it in the form of light (a stream of photons). The mechanism of gamma quantum energy absorption within the scintillator may be of dual nature, occurring either by means of the Compton's effect or by means of the photoelectric phenomenon, with only the photoelectric phenomenon being taken into account in calculations carried out by current PET scanners. Thus, it is assumed that the number of photons generated in the scintillator material is proportional to the energy of gamma quanta deposited within the scintillator.
When two annihilation gamma quanta are detected by a pair of detectors at a time interval not larger than several nanoseconds, i.e. in coincidence, the position of annihilation point along the line of response may be determined, i.e. along the line connecting the detector centers or the points within the scintillator strips where the energy of the gamma quanta was deposited. The coordinates of annihilation place are obtained from the difference in times of arrival of two gamma quanta to the detectors located at both ends of the LOR. In the prior art literature, this technique is referred to as the time of flight (TOF) technique and the PET scanners utilizing time measurements are referred to as TOF-PET scanners. This technique requires that the scintillator has a time resolution of a few hundred picoseconds.
Currently, the state of the art methods of determining the places of interactions of the gamma quanta in positron emission tomography are based on the measurements of charges of signals generated in vacuum tube photomultipliers, silicon photomultipliers, or avalanche diodes optically connected to inorganic crystals notched into smaller elements. Position of the gamma quantum reaction is determined with the accuracy of the smaller crystal element size on the basis of the differences in changes of the signals from different converters optically connected to the same crystal. In the state of the art PET scanners, reconstruction of the set of LOR and TOF data is based on the relationships between charges and times of signals recorded for a particular event without reference to external signals.
In the signal time determination methods used in the state of the art, changes in shapes and amplitudes of signals depending on the place of ionization and the quantity of energy constitute a limitation in temporal resolutions that can be achieved using the technique. The larger the scintillator, the larger the variations in signal shapes and amplitudes.
For the above, reasons, temporal resolutions of less than 100 ps are unattainable in the state of the art for large scintillator blocks. Temporal resolution also translates on the resolution of ionization position determination. In case of polymer scintillators (preferred due to their low price), amplitudes of signals generated by the gamma quanta, including annihilation gamma quanta used in positron emission tomography, are characterized by continuous distribution resulting from interactions between gamma quanta and electrons occurring mostly via the Compton effect with a negligibly low probability of a photoelectric effect. As a consequence, signal amplitudes in polymer scintillators may change even if they originate from the same position.
As shown by the shortcomings of the state of the art signal analysis techniques, there is a need to significantly improve temporal and spatial resolution in the detectors used in medical diagnostic techniques that require recording of ionizing radiation. The need to improve resolution is particularly high in large-sized detectors. Examples of PET detectors making use of large polymeric scintillators were described in patent application WO 2011/008119 as well as in patent application WO 2011/008118. Solutions described in these applications are based on the measurements of the times of light pulses arrival to the detector edges. Light pulses are converted into electric pulses by means of photomultipliers. The shape (temporal distribution of photons) and the amplitude of the light pulse reaching the photomultiplier changes depending on the distance between the photomultiplier and the pulse origin place. In addition and independently of the ionization place, the amplitude of the signal changes with the energy deposited within the detector. As a consequence, due to variations in signal shapes and amplitudes, it is impossible to achieve good temporal resolution using either leading edge or constant fraction discriminators of the current state of the art due to the time walk effect and the pulse shape change effect observed in large-size scintillators.
PET detectors require time and energy calibration that is carried out using radioactive isotopes such as 22Na or 68Ge, placed in precisely defined positions within a PET scanner, for example in the geometric center of the scanner or used as a mobile radiation sources rotating around the scintillation chamber, facilitating relative synchronization of all detection elements.
Methods for energy calibration of the detection systems in TOF-PET scanners are known in patent literature.
U.S. Pat. No. 7,414,246 and U.S. Pat. No. 7,820,975 disclose methods of timing calibration of detectors in TOF-PET scanners wherein sodium isotope placed in a metal or plastic shield is used as the radiation source and the annihilation quanta scattered at the shield are used for determination of the relative delays of individual detection modules within the PET scanner.
The U.S. Pat. No. 7,557,350 discloses a method of temporal synchronization of TOF-PET detectors wherein several radioactive sources are used at the same time, facilitating timing calibration of TOF-PET scanners to be carried out even while acquiring object images. The gamma quanta originating from the calibration sources according to the disclosed method are discriminated on the basis of the known positions of the energy sources as well as the timing information from detectors, which additionally permits discarding these events when reconstructing tomographic images.
The U.S. Pat. No. 5,272,343 discloses a method for synchronization of PET detectors making use of the orbiting of a radiation source. The annihilation gamma quanta from the moving radioactive source facilitate synchronization of PET scanned detector pairs by making use of the fact that in the case of the radiation source orbiting around the PET scanner axis, the difference in times of recording these quanta by two detectors located opposite each other is constant regardless of the position of the source within the scanner.
However, the described methods for calibration of detectors used in TOF-PET scanners do not permit calibration of detectors while acquiring object scans without the risk of exposing the object to an additional dose of radiation emitted by the radioactive sources used for calibration. In addition, the use of radioactive sources for synchronization of TOF-PET detectors requires additional equipment, trainings of personnel operating TOF-PET scanners and replacement of the sources as they decline in activity (for example, the half-life of 68Ge decay is about 270 days), thus increasing the imaging costs. Currently, calibration of TOF-PET detection systems are usually performed once a day before acquiring object scans so as to avoid the object's exposure to additional radiation; however, this method does not allow consideration of the changing ambient conditions, i.e. fluctuations in temperature or voltage that may affect temporal or energy properties of TOF-PET detectors while shortening the time devoted to the scanning of individual objects. In addition, the aforementioned methods for TOF-PET calibration are not convenient when using long TOF-PET detectors as employed in strip or matrix TOF-PET scanners disclosed in patent applications WO 2011/008119 and WO 2011/008118, where polymeric scintillation material has been used in the form of long strips or plates connected to photomultiplier systems for recording the annihilation quanta.
It would be expedient to develop a method for calibration of TOF-PET detectors and monitoring of the quality of the detection system that would facilitate continuous monitoring of the detection system and simultaneous calibration of TOF-PET detectors while performing imaging scans and not requiring the use of additional radiation sources.